Integrated lancing and measurement device

ABSTRACT

An integrated lancing and measurement device is provided comprising a sensor designed to determine the amount and/or concentration of analyte in a biological fluid having a volume of less than about 1 μL. A piercing member is adapted to pierce and retract from a site on the patient to cause the fluid to flow therefrom, and the sensor is positioned adjacent to the site on the patient so as to receive the fluid flowing from the site to generate an electrical signal indicative of the concentration of the analyte in the fluid. The sensor is comprised of a working electrode comprising an analyte-responsive enzyme and a redox mediator, and a counter electrode. An analyte monitor is operatively connected to the sensor and adapted to measure the signal generated by the sensor. Also provided are analyte measuring methods that optionally employ the integrated lancing and measurement device.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application is a Continuation of U.S. Ser. No. 10/629,348, filedJul. 28, 2003, published as US2004/0055898, now pending, which is acontinuation of U.S. Ser. No. 09/714,360, filed Nov. 15, 2000, now U.S.Pat. No. 6,607,658, which is a continuation of U.S. Ser. No. 09/544,593,filed Apr. 6, 2000, now U.S. Pat. No. 6,551,494, which is a continuationof U.S. Ser. No. 09/326,235, filed Jun. 4, 1999, now U.S. Pat. No.6,120,676, which is a continuation of U.S. Ser. No. 08/795,767, filedFeb. 6, 1997, now abandoned, which applications are incorporated byreference herein.

TECHNICAL FIELD

This invention relates to analytical sensors for the detection ofbioanalytes in a small volume sample.

BACKGROUND OF THE INVENTION

Analytical sensors are useful in chemistry and medicine to determine thepresence and concentration of a biological analyte. Such sensors areneeded, for example, to monitor glucose in diabetic patients and lactateduring critical care events.

Currently available technology measures bioanalytes in relatively largesample volumes, e.g., generally requiring 3 microliters or more of bloodor other biological fluid. These fluid samples are obtained from apatient, for example, using a needle and syringe, or by lancing aportion of the skin such as the fingertip and “milking” the area toobtain a useful sample volume. These procedures are inconvenient for thepatient, and often painful, particularly when frequent samples arerequired. Less painful methods for obtaining a sample are known such aslancing the arm or thigh, which have a lower nerve ending density.However, lancing the body in the preferred regions typically producessubmicroliter samples of blood, because these regions are not heavilysupplied with near-surface capillary vessels.

It would therefore be desirable and very useful to develop a relativelypainless, easy to use blood analyte sensor, capable of performing anaccurate and sensitive analysis of the concentration of analytes in asmall volume of sample.

The sensors of the present invention provide a method for the detectionand quantification of an analyte in submicroliter samples. In general,the invention includes a method and sensor for analysis of an analyte ina small volume of sample, preferably by coulometry. A biosensor of theinvention utilizes a non-leachable redox mediator, preferably anair-oxidizable redox mediator, and preferably immobilized on a workingelectrode. The biosensor also includes a sample chamber to hold thesample in electrolytic contact with the working electrode. In apreferred embodiment, the working electrode faces a counter electrode,forming a measurement zone within the sample chamber, between the twoelectrodes, that is sized to contain less than about 1 μL of sample,preferably less than about 0.5 μL, more preferably less than about 0.2μL, and most preferably less than about 0.1 μL of sample. A sorbentmaterial is optionally positioned in the sample chamber and measurementzone to reduce the volume of sample needed to fill the sample chamberand measurement zone.

In one embodiment of the invention, a biosensor is provided whichcombines the efficiency of coulometric electrochemical sensing with anon-leachable redox mediator to accurately and efficiently measure abioanalyte in a submicroliter volume of sample. The preferred sensorincludes an electrode, a non-leachable redox mediator on the electrode,a sample chamber for holding the sample in electrical contact with theelectrode and, preferably, sorbent material disposed within the samplechamber to reduce the volume of the chamber. The sample chamber,together with any sorbent material, is sized to provide for analysis ofa sample volume that is typically less than about 1 μL, preferably lessthan about 0.5 μL, more preferably less than about 0.2 μL, and mostpreferably less than about 0.1 μL.

One embodiment of the invention includes a method for determining theconcentration of an analyte in a sample by, first, contacting the samplewith an electrochemical sensor and then determining the concentration ofthe analyte. The electrochemical sensor includes a facing electrode pairwith a working electrode and a counter electrode and a sample chamber,including a measurement zone, positioned between the two electrodes. Themeasurement zone is sized to contain less than about 1 μL of sample.

The invention also includes an electrochemical sensor with two or morefacing electrode pairs. Each electrode pair has a working electrode, acounter electrode, and a measurement zone between the two electrodes,the measurement zone being sized to hold less than about 1 μL of sample.In addition, the sensor also includes non-leachable redox mediator onthe working electrode of at least one of the electrode pairs.

One aspect of the invention is a method of determining the concentrationof an analyte in a sample by contacting the sample with anelectrochemical sensor and determining the concentration of the analyteby coulometry. The electrochemical sensor includes an electrode pairwith a working electrode and a counter electrode. The sensor alsoincludes a sample chamber for holding a sample in electrolytic contactwith the working electrode. Within the sample chamber is sorbentmaterial to reduce the volume sample needed to fill the sample chamberso that the sample chamber is sized to contain less than about 1 μL ofsample.

The invention also includes a sensor and a method for the determinationof the concentration of an analyte in a sample having a volume of lessthan about 1 μL. The sensor has a support and an air-oxidizable redoxmediator coated on the support. At least 90% of the air-oxidizable redoxmediator is in an oxidized state prior to introduction of a sample. Themethod includes contacting the sample with the sensor and correlatingthe concentration of the analyte in the sample to a change in oxidationstate of the redox mediator in the presence of the sample. The sensorand method of this aspect of the invention are directed to, but notlimited to, electrochemical and optical sensors.

A further aspect of the invention is an integrated sample acquisitionand analyte measurement device which includes a sample acquisition meansfor producing a patient sample as well as a sensor of the invention formeasuring analyte in the sample. The device is used for measuringanalyte in a patient sample by, first, contacting the patient with thedevice and then determining the concentration of the analyte, preferablyby coulometry.

Another aspect of the invention is a method for determining theconcentration of an analyte in the sample with reduced error bycontacting the sample with an electrochemical sensor that includes afirst and a second electrode pair. Each electrode pair has a workingelectrode and a sample chamber for holding the sample in electrolyticcontact with the working electrode, the sample chamber being sized tocontain less than about 1 μL of sample. The first electrode pair alsohas a non-leachable redox mediator and non-leachable enzyme on theworking electrode. The second electrode pair has a non-leachable redoxmediator in the absence of enzyme on the working electrode. The methodfurther includes the step of measuring substantially simultaneously, andat two or more times, a first current generated at the first electrodepair and a second current generated at the second electrode pair. Themeasured first currents and second currents are independently integratedto give a first charge and a second charge, respectively. The secondcharge is subtracted from the first charge to give a noise-reducedcharge which is then correlated to the concentration of analyte in thesample. This method can be used to remove errors arising frominterferents or the mixed oxidation state of the redox mediator prior tointroduction of the sample.

Another method of the invention for the determination of theconcentration of an analyte in a sample includes the step of providingan electrochemical sensor which has one or more facing electrode pairs,each pair having a working and a counter electrode and a measurementzone between the working and counter electrodes, the measurement zonesof the one or more electrode pairs having approximately equal volumes ofless than about 1 μL. The sensor also includes redox mediator on theworking electrode of at least one of the electrode pairs. The methodfurther includes measuring a capacitance of one of the electrode pairsand calculating the volume of the measurement zone of that electrodepair from the capacitance measurement. In addition, the sensor isbrought into contact with the sample and the concentration of analyte inthe sample is determined by coulometry.

A further aspect of the invention is a method of storing and packagingan analytical sensor which includes packaging the sensor in anatmosphere containing molecular oxygen. The sensor of this aspect of theinvention includes air-oxidizable redox mediator.

One embodiment of the invention is a method of determining theconcentration of an analyte in a sample by contacting the sample with anelectrochemical sensor, electrolyzing less than about 1 μL of sample,and determining the concentration of the analyte by coulometry. Thesensor of this embodiment of the invention includes a working electrodeand non-leachable redox mediator on the working electrode. The molaramount of non-leachable redox mediator in the reduced form prior tointroduction of the sample into the sensor is less than, on astoichiometric basis, 5% of the expected molar amount of analyte to beelectrolyzed.

Another method for determining the concentration of an analyte in asample includes contacting the sample with an electrochemical sensorwhich has a working electrode, a counter electrode, and a measurementzone bounded on at least two sides by the two electrodes. Themeasurement zone is sized to contain less than about 1 μL of sample. Theconcentration of analyte in the sample is then determined by coulometry.

These and various other features which characterize the invention arepointed out with particularity in the attached claims. For a betterunderstanding of the invention, its advantages, and objectives obtainedby its use, reference should be made to the drawings and to theaccompanying description, in which there is illustrated and describedpreferred embodiments of the invention.

BRIEF DESCRIPTION OF THE DRAWINGS

Referring now to the drawings, wherein like reference numerals andletters indicate corresponding structure throughout the several views:

FIG. 1 is a schematic view of a first embodiment of an electrochemicalsensor in accordance with the principles of the present invention havinga working electrode and a counter electrode facing each other;

FIG. 2 is a schematic view of a second embodiment of an electrochemicalsensor in accordance with the principles of the present invention havinga working electrode and a counter electrode in a coplanar configuration;

FIG. 3 is a schematic view of a third embodiment of an electrochemicalsensor in accordance with the principles of the present invention havinga working electrode and a counter electrode facing each other and havingan extended sample chamber;

FIG. 4 is a not-to-scale side-sectional drawing of a portion of thesensor of FIG. 1 or 3 showing the relative positions of the redoxmediator, the sample chamber, and the electrodes;

FIG. 5 is a top view of an embodiment of a multiple electrode sensor inaccordance with the principles of the present invention;

FIG. 6 is a perspective view of an embodiment of an analyte measurementdevice in accordance with the principles of the present invention havinga sample acquisition means and the sensor of FIG. 4;

FIG. 7 is a graph of the charge required to electrooxidize a knownquantity of glucose in an electrolyte buffered solution (filled circles)or serum solution (open circles) using the sensor of FIG. 1 with glucoseoxidase as the second electron transfer agent;

FIG. 8 is a graph of the average glucose concentrations for the data ofFIG. 7 (buffered solutions only) with calibration curves calculated tofit the averages; a linear calibration curve was calculated for the10-20 mM concentrations and a second order polynomial calibration curvewas calculated for the 0-10 mM concentrations;

FIG. 9 is a Clarke-type clinical grid analyzing the clinical relevanceof the glucose measurements of FIG. 7; and

FIG. 10 is a graph of the charge required to electrooxidize a knownquantity of glucose in an electrolyte buffered solution using the sensorof FIG. 1 with glucose dehydrogenase as the second electron transferagent.

DETAILED DESCRIPTION OF THE INVENTION

When used herein, the following definitions define the stated term:

An “air-oxidizable mediator” is a redox mediator that is oxidized byair, preferably so that at least 90% of the mediator is in an oxidizedstate upon storage in air within a useful period of time, e.g., onemonth or less, and, preferably, one week or less, and, more preferably,one day or less.

A “biological fluid” is any body fluid in which the analyte can bemeasured, for example, blood, interstitial fluid, dermal fluid, sweat,and tears.

The term “blood” in the context of the invention includes whole bloodand its cell-free components, namely, plasma and serum.

“Coulometry” is the determination of charge passed or projected to passduring complete or nearly complete electrolysis of the analyte, eitherdirectly on the electrode or through one or more electron transferagents. The charge is determined by measurement of charge passed duringpartial or nearly complete electrolysis of the analyte or, more often,by multiple measurements during the electrolysis of a decaying currentand elapsed time. The decaying current results from the decline in theconcentration of the electrolyzed species caused by the electrolysis.

A “counter electrode” refers to an electrode paired with the workingelectrode, through which passes an electrochemical current equal inmagnitude and opposite in sign to the current passed through the workingelectrode. In the context of the invention, the term “counter electrode”is meant to include counter electrodes which also function as referenceelectrodes (i.e. a counter/reference electrode).

An “electrochemical sensor” is a device configured to detect thepresence and/or measure the concentration of an analyte viaelectrochemical oxidation and reduction reactions on the sensor. Thesereactions are transduced to an electrical signal that can be correlatedto an amount or concentration of analyte.

“Electrolysis” is the electrooxidation or electroreduction of a compoundeither directly at an electrode or via one or more electron transferagents.

The term “facing electrodes” refers to a configuration of the workingand counter electrodes in which the working surface of the workingelectrode is disposed in approximate opposition to a surface of thecounter electrode and where the distance between the working and counterelectrodes is less than the width of the working surface of the workingelectrode.

A compound is “immobilized” on a surface when it is entrapped on orchemically bound to the surface.

The “measurement zone” is defined herein as a region of the samplechamber sized to contain only that portion of the sample that is to beinterrogated during the analyte assay.

A “non-leachable” or “non-releasable” compound is a compound which doesnot substantially diffuse away from the working surface of the workingelectrode for the duration of the analyte assay.

A “redox mediator” is an electron transfer agent for carrying electronsbetween the analyte and the working electrode, either directly, or via asecond electron transfer agent.

A “second electron transfer agent” is a molecule which carries electronsbetween the redox mediator and the analyte.

“Sorbent material” is material which wicks, retains, or is wetted by afluid sample in its void volume and which does not substantially preventdiffusion of the analyte to the electrode.

A “working electrode” is an electrode at which analyte iselectrooxidized or electroreduced with or without the agency of a redoxmediator.

A “working surface” is that portion of the working electrode which iscoated with redox mediator and configured for exposure to sample.

The small volume, in vitro analyte sensors of the present invention aredesigned to measure the concentration of an analyte in a portion of asample having a volume less than about 1 μL, preferably less than about0.5 μL, more preferably less than about 0.2 μL, and most preferably lessthan about 0.1 μL. The analyte of interest is typically provided in asolution or biological fluid, such as blood or serum. Referring to theDrawings in general and FIGS. 1-4 in particular, a small volume, invitro electrochemical sensor 20 of the invention generally includes aworking electrode 22, a counter (or counter/reference) electrode 24, anda sample chamber 26 (see FIG. 4). The sample chamber 26 is configured sothat when a sample is provided in the chamber the sample is inelectrolytic contact with both the working electrode 22 and the counterelectrode 24. This allows electrical current to flow between theelectrodes to effect the electrolysis (electrooxidation orelectroreduction) of the analyte.

Working Electrode

The working electrode 22 may be formed from a molded carbon fibercomposite or it may consist of an inert non-conducting base material,such as polyester, upon which a suitable conducting layer is deposited.The conducting layer should have relatively low electrical resistanceand should be electrochemically inert over the potential range of thesensor during operation. Suitable conductors include gold, carbon,platinum, ruthenium dioxide and palladium, as well as othernon-corroding materials known to those skilled in the art. The electrodeand/or conducting layers are deposited on the surface of the inertmaterial by methods such as vapor deposition or printing.

A tab 23 may be provided on the end of the working electrode 22 for easyconnection of the electrode to external electronics (not shown) such asa voltage source or current measuring equipment. Other known methods orstructures may be used to connect the working electrode 22 to theexternal electronics.

Sensing Layer and Redox Mediator

A sensing layer 32 containing a non-leachable (i.e., non-releasable)redox mediator is disposed on a portion of the working electrode 22.Preferably, there is little or no leaching of the redox mediator awayfrom the working electrode 22 into the sample during the measurementperiod, which is typically less than about 5 minutes. More preferably,the redox mediators of the present invention are bound or otherwiseimmobilized on the working electrode 22 to prevent undesirable leachingof the mediator into the sample. A diffusing or leachable (i.e.,releasable) redox mediator is not desirable when the working and counterelectrodes are close together (i.e., when the electrodes are separatedby less than about 1 mm), because a large background signal is typicallyproduced as the unbound mediator shuttles electrons between the workingand counter electrodes, rather than between the analyte and the workingelectrode. This and other problems have hindered the development of lowresistance cells and increased the minimum sample size required fordetermination of analyte concentration.

Application of sensing layer 32 on working electrode 22 creates aworking surface on that electrode. In general, the working surface isthat portion of the working electrode 22 coated with mediator and ableto contact a fluid sample. If a portion of the sensing layer 32 iscovered by a dielectric or other material, then the working surface willonly be that portion of the electrode covered by redox mediator andexposed for contact with the sample.

The redox mediator mediates a current between the working electrode 22and the analyte and enables the electrochemical analysis of moleculeswhich are not suited for direct electrochemical reaction on anelectrode. The mediator functions as an electron transfer agent betweenthe electrode and the analyte.

Almost any organic or organometallic redox species can be used as aredox mediator. In general, the preferred redox mediators are rapidlyreducible and oxidizable molecules having redox potentials a few hundredmillivolts above or below that of the standard calomel electrode (SCE),and typically not more reducing than about −100 mV and not moreoxidizing than about +400 mV versus SCE. Examples of organic redoxspecies are quinones and quinhydrones and species that in their oxidizedstate have quinoid structures, such as Nile blue and indophenol.Unfortunately, some quinones and partially oxidized quinhydrones reactwith functional groups of proteins such as the thiol groups of cysteine,the amine groups of lysine and arginine, and the phenolic groups oftyrosine which may render those redox species unsuitable for some of thesensors of the present invention, e.g., sensors that will be used tomeasure analyte in biological fluids such as blood.

In general, mediators suitable for use in the invention have structureswhich prevent or substantially reduce the diffusional loss of redoxspecies during the period of time that the sample is being analyzed. Thepreferred redox mediators include a redox species bound to a polymerwhich can in turn be immobilized on the working electrode. Useful redoxmediators and methods for producing them are described in U.S. Pat. Nos.5,264,104; 5,356,786; 5,262,035; and 5,320,725, herein incorporated byreference. Although, any organic or organometallic redox species can bebound to a polymer and used as a redox mediator, the preferred redoxspecies is a transition metal compound or complex. The preferredtransition metal compounds or complexes include osmium, ruthenium, iron,and cobalt compounds or complexes. The most preferred are osmiumcompounds and complexes.

One type of non-releasable polymeric redox mediator contains a redoxspecies covalently bound in a polymeric composition. An example of thistype of mediator is poly(vinylferrocene).

Alternatively, a suitable non-releasable redox mediator contains anionically-bound redox species. Typically, these mediators include acharged polymer coupled to an oppositely charged redox species. Examplesof this type of mediator include a negatively charged polymer such asNafion® (Dupont) coupled to a positively charged redox species such asan osmium or ruthenium polypyridyl cation. Another example of anionically-bound mediator is a positively charged polymer such asquaternized poly(4-vinyl pyridine) or poly(1-vinyl imidazole) coupled toa negatively charged redox species such as ferricyanide or ferrocyanide.

In another embodiment of the invention, the suitable non-releasableredox mediators include a redox species coordinatively bound to thepolymer. For example, the mediator may be formed by coordination of anosmium or cobalt 2,2′-bipyridyl complex to poly(1-vinyl imidazole) orpoly(4-vinyl pyridine).

The preferred redox mediators are osmium transition metal complexes withone or more ligands having a nitrogen-containing heterocycle such as2,2′-bipyridine, 1,10-phenanthroline or derivatives thereof.Furthermore, the preferred redox mediators also have one or morepolymeric ligands having at least one nitrogen-containing heterocycle,such as pyridine, imidazole, or derivatives thereof. These preferredmediators exchange electrons rapidly between each other and theelectrodes so that the complex can be rapidly oxidized and reduced.

In particular, it has been determined that osmium cations complexed withtwo ligands containing 2,2′-bipyridine, 1,10-phenanthroline, orderivatives thereof, the two ligands not necessarily being the same, andfurther complexed with a polymer having pyridine or imidazole functionalgroups form particularly useful redox mediators in the small volumesensors of the present invention. Preferred derivatives of2,2′-bipyridine for complexation with the osmium cation are4,4′-dimethyl-2,2′-bipyridine and mono-, di-, andpolyalkoxy-2,2′-bipyridines, such as 4,4′-dimethoxy-2,2′-bipyridine,where the carbon to oxygen ratio of the alkoxy groups is sufficient toretain solubility of the transition metal complex in water. Preferredderivatives of 1,10-phenanthroline for complexation with the osmiumcation are 4,7-dimethyl-1,10-phenanthroline and mono-, di-, andpolyalkoxy-1,10-phenanthrolines, such as4,7-dimethoxy-1,10-phenanthroline, where the carbon to oxygen ratio ofthe alkoxy groups is sufficient to retain solubility of the transitionmetal complex in water. Preferred polymers for complexation with theosmium cation include poly(1-vinyl imidazole), e.g., PVI, andpoly(4-vinyl pyridine), e.g., PVP, either alone or with a copolymer.Most preferred are redox mediators with osmium complexed withpoly(1-vinyl imidazole) alone or with a copolymer.

The preferred redox mediators have a redox potential between about −150mV to about +400 mV versus the standard calomel electrode (SCE).Preferably, the potential of the redox mediator is between about −100 mVand +100 mV and more preferably, the potential is between about −50 mVand +50 mV. The most preferred redox mediators have osmium redox centersand a redox potential more negative than +100 mV versus SCE, morepreferably the redox potential is more negative than +50 mV versus SCE,and most preferably is near −50 mV versus SCE.

It is also preferred that the redox mediators of the inventive sensorsbe air-oxidizable. This means that the redox mediator is oxidized byair, preferably so that at least 90% of the mediator is in an oxidizedstate prior to introduction of sample into the sensor. Air-oxidizableredox mediators include osmium cations complexed with two mono-, di-, orpolyalkoxy-2,2′-bipyridine or mono-, di-, orpolyalkoxy-1,10-phenanthroline ligands, the two ligands not necessarilybeing the same, and further complexed with polymers having pyridine andimidazole functional groups. In particular,Os[4,4′-dimethoxy-2,2′-bipyridine]₂Cl^(+/+2) complexed with poly(4-vinylpyridine) or poly(1-vinyl imidazole) attains approximately 90% or moreoxidation in air.

In a preferred embodiment of the invention, the sensing layer 32includes a second electron transfer agent which is capable oftransferring electrons to or from the redox mediator and the analyte.One example of a suitable second electron transfer agent is an enzymewhich catalyzes a reaction of the analyte. For example, a glucoseoxidase or glucose dehydrogenase, such as pyrroloquinoline quinoneglucose dehydrogenase (PQQ), is used when the analyte is glucose. Alactate oxidase fills this role when the analyte is lactate. Theseenzymes catalyze the electrolysis of an analyte by transferringelectrons between the analyte and the electrode via the redox mediator.Preferably, the second electron transfer agent is non-leachable, andmore preferably immobilized on the electrode, to prevent unwantedleaching of the agent into the sample. This is accomplished, forexample, by crosslinking the second electron transfer agent with theredox mediator, thereby providing a sensing layer with non-leachablecomponents.

To prevent electrochemical reactions from occurring on portions of theworking electrode not coated by the mediator, a dielectric 40 may bedeposited on the electrode over, under, or surrounding the region withthe bound redox mediator, as shown in FIG. 4. Suitable dielectricmaterials include waxes and non-conducting organic polymers such aspolyethylene. Dielectric 40 may also cover a portion of the redoxmediator on the electrode. The covered portion of the mediator will notcontact the sample, and, therefore, will not be a part of theelectrode's working surface.

Counter Electrode

Counter electrode 24 may be constructed in a manner similar to workingelectrode 22. Counter electrode 24 may also be a counter/referenceelectrode. Alternatively, a separate reference electrode may be providedin contact with the sample chamber. Suitable materials for thecounter/reference or reference electrode include Ag/AgCl printed on anon-conducting base material or silver chloride on a silver metal base.If the counter electrode is not a reference electrode, the samematerials and methods may be used to make the counter electrode as areavailable for constructing the working electrode 22, however, no redoxmediator is immobilized on the counter or counter/reference electrode24. A tab 25 may be provided on the electrode for convenient connectionto the external electronics (not shown), such as a coulometer or othermeasuring device.

In one embodiment of the invention, working electrode 22 and counterelectrode 24 are disposed opposite to and facing each other to form afacing electrode pair as depicted in FIGS. 1 and 3. In this preferredconfiguration, the sample chamber 26 is typically disposed between thetwo electrodes. For this facing electrode configuration, it is preferredthat the electrodes are separated by a distance of less than about 0.2mm, preferably less than 0.1 mm, and most preferably less than 0.05 mm.

The electrodes need not be directly opposing each other, they may beslightly offset. Furthermore, the two electrodes need not be the samesize. Preferably, the counter electrode 24 is at least as large as theworking surface of the working electrode 22. The counter electrode 22can also be formed with tines in a comb shape. Other configuration ofboth the counter electrode and working electrode are within the scope ofthe invention. However, the separation distance between any portion ofthe working electrode and some portion of the counter electrodepreferably does not exceed the limits specified hereinabove.

In another embodiment of the invention, the two electrodes 22, 24 arecoplanar as shown in FIG. 2. In this case, the sample chamber 26 is incontact with both electrodes and is bounded on the side opposite theelectrodes by a non-conducting inert base 30. Suitable materials for theinert base include non-conducting materials such as polyester.

Other configurations of the inventive sensors are also possible. Forexample, the two electrodes may be formed on surfaces that make an angleto each other. One such configuration would have the electrodes onsurfaces that form a right angle. Another possible configuration has theelectrodes on a curved surface such as the interior of a tube. Theworking and counter electrodes may be arranged so that they face eachother from opposite sides of the tube. This is another example of afacing electrode pair. Alternatively, the electrodes may be placed neareach other on the tube wall (e.g., one on top of the other orside-by-side).

In any configuration, the two electrodes must be configured so that theydo not make direct electrical contact with each other, to preventshorting of the electrochemical sensor. This may be difficult to avoidwhen the facing electrodes having a short (less than about 100 μm)distance between them.

A spacer 28 can be used to keep the electrodes apart when the electrodesface each other as depicted in FIGS. 1 and 3. The spacer is typicallyconstructed from an inert non-conducting material such as polyester,Mylar™, Kevlar™ or any other strong, thin polymer film, or,alternatively, a thin polymer film such as a Teflon™ film, chosen forits chemical inertness. In addition to preventing contact between theelectrodes, the spacer 28 often functions as a portion of the boundaryfor the sample chamber 26 as shown in FIGS. 1-4.

Sample Chamber

The sample chamber 26 is typically defined by a combination of theelectrodes 22, 24, an inert base 30, and a spacer 28 as shown in FIGS.1-4. A measurement zone is contained within this sample chamber and isthe region of the sample chamber that contains only that portion of thesample that is interrogated during the analyte assay. In the embodimentof the invention illustrated in FIGS. 1 and 2, sample chamber 26 is thespace between the two electrodes 22, 24 and/or the inert base 30. Inthis embodiment, the sample chamber has a volume that is preferably lessthan about 1 μL, more preferably less than about 0.5 μL, and mostpreferably less than about 0.2 μL. In the embodiment of the inventiondepicted in FIGS. 1 and 2, the measurement zone has a volume that isapproximately equal to the volume of the sample chamber.

In another embodiment of the invention, shown in FIG. 3, sample chamber26 includes much more space than the region proximate electrodes 22, 24.This configuration makes it possible to provide multiple electrodes incontact with one or more sample chambers, as shown in FIG. 5. In thisembodiment, sample chamber 26 is preferably sized to contain a volume ofless than about 1 μL, more preferably less than about 0.5 μL, and mostpreferably less than about 0.2 μL. The measurement zone (i.e., theregion containing the volume of sample to be interrogated) is generallysized to contain a volume of sample of less than about 1 μL, preferablyless than about 0.5 μL, more preferably less than about 0.2 μL, and mostpreferably less than about 0.1 μL. One particularly useful configurationof this embodiment positions working electrode 22 and counter electrode24 facing each other, as shown in FIG. 3. In this embodiment, themeasurement zone, corresponding to the region containing the portion ofthe sample which will be interrogated, is the portion of sample chamber26 bounded by the working surface of the working electrode and disposedbetween the two facing electrodes. When the surface of the workingelectrode is not entirely covered by redox mediator, the measurementzone is the space between the two facing electrodes that has a surfacearea corresponding to the working surface (i.e., redox mediator-coveredsurface) of working electrode 22 and a thickness corresponding to theseparation distance between working electrode 22 and counter electrode24.

In both of the embodiments discussed above, the thickness of the samplechamber and of the measurement zone correspond typically to thethickness of spacer 28 (e.g., the distance between the electrodes inFIGS. 1 and 3, or the distance between the electrodes and the inert basein FIG. 2). Preferably, this thickness is small to promote rapidelectrolysis of the analyte, as more of the sample will be in contactwith the electrode surface for a given sample volume. In addition, athin sample chamber helps to reduce errors from diffusion of analyteinto the measurement zone from other portions of the sample chamberduring the analyte assay, because diffusion time is long relative to themeasurement time. Typically, the thickness of the sample chamber is lessthan about 0.2 mm. Preferably, the thickness of the sample chamber isless than about 0.1 mm and, more preferably, the thickness of the samplechamber is about 0.05 mm or less.

The sample chamber may be empty before the sample is placed in thechamber. Alternatively, the sample chamber may include a sorbentmaterial 34 to sorb and hold a fluid sample during the measurementprocess. Suitable sorbent materials include polyester, nylon, cellulose,and cellulose derivatives such as nitrocellulose. The sorbent materialfacilitates the uptake of small volume samples by a wicking action whichmay complement or, preferably, replace any capillary action of thesample chamber.

One of the most important functions of the sorbent material is to reducethe volume of fluid needed to fill the sample chamber and correspondingmeasurement zone of the sensor. The actual volume of sample within themeasurement zone is partially determined by the amount of void spacewithin the sorbent material. Typically, suitable sorbents consist ofabout 5% to about 50% void space. Preferably, the sorbent materialconsists of about 10% to about 25% void space.

The displacement of fluid by the sorbent material is advantageous. Byaddition of a sorbent, less sample is needed to fill sample chamber 26.This reduces the volume of sample that is required to obtain ameasurement and also reduces the time required to electrolyze thesample.

The sorbent material 34 may include a tab 33 which is made of the samematerial as the sorbent and which extends from the sensor, or from anopening in the sensor, so that a sample may be brought into contact withtab 33, sorbed by the tab, and conveyed into the sample chamber 26 bythe wicking action of the sorbent material 34. This provides a preferredmethod for directing the sample into the sample chamber 26. For example,the sensor may be brought into contact with a region of an animal(including human) that has been pierced with a lancet to draw blood. Theblood is brought in contact with tab 33 and drawn into sample chamber 26by the wicking action of the sorbent 34. The direct transfer of thesample to the sensor is especially important when the sample is verysmall, such as when the lancet is used to pierce a portion of the animalthat is not heavily supplied with near-surface capillary vessels andfurnishes a blood sample volume of less than 1 μL.

Methods other than the wicking action of a sorbent may be used totransport the sample into the sample chamber or measurement zone.Examples of such means for transport include the application of pressureon a sample to push it into the sample chamber, the creation of a vacuumby a pump or other vacuum-producing means in the sample chamber to pullthe sample into the chamber, capillary action due to interfacial tensionof the sample with the walls of a thin sample chamber, as well as thewicking action of a sorbent material.

The sensor can also be used in conjunction with a flowing sample stream.In this configuration, the sample stream is made to flow through asample chamber. The flow is stopped periodically and the concentrationof the analyte is determined by electrochemical method, such ascoulometry. After the measurement, the flow is resumed, thereby removingthe sample from the sensor. Alternatively, sample may flow through thechamber at a very slow rate, such that all of the analyte iselectrolyzed in transit, yielding a current dependent only upon analyteconcentration and flow rate.

The entire sensor assembly is held firmly together to ensure that thesample remains in contact with the electrodes and that the samplechamber and measurement zone maintain the same volume. This is animportant consideration in the coulometric analysis of a sample, wheremeasurement of a defined sample volume is needed. One method of holdingthe sensor together is depicted in FIGS. 1 and 2. Two plates 38 areprovided at opposite ends of the sensor. These plates are typicallyconstructed of non-conducting materials such as plastics. The plates aredesigned so that they can be held together with the sensor between thetwo plates. Suitable holding devices include adhesives, clamps, nuts andbolts, screws, and the like.

Integrated Sample Acquisition and Analyte Measurement Device

In a preferred embodiment of the invention, an analyte measurementdevice 52 constructed according to the principles of the presentinvention includes a sensor 20, as described hereinabove, combined witha sample acquisition means 50 to provide an integrated sampling andmeasurement device. The sample acquisition means 50 illustrated in FIG.6, includes, for example, a skin piercing member 54, such as a lancet,attached to a resilient deflectable strip 56 (or other similar device,such as a spring) which may be pushed to inject the lancet into apatient's skin to cause blood flow.

The resilient strip 56 is then released and the skin piercing member 54retracts. Blood flowing from the area of skin pierced by member 54 canthen be transported, for example, by the wicking action of sorbentmaterial 34, into sensor 20 for analysis of the analyte. The analytemeasurement device 52 may then be placed in a reader, not shown, whichconnects a coulometer or other electrochemical analysis equipment to theelectrode tabs 23, 25 to determine the concentration of the analyte byelectroanalytical means.

Operation of the Sensor

An electrochemical sensor of the invention is operated in the followingmanner. A potential is applied across the working and counterelectrodes. The magnitude of the required potential is dependent on theredox mediator. The potential at an electrode where the analyte iselectrolyzed is typically large enough to drive the electrochemicalreaction to or near completion, but the magnitude of the potential is,preferably, not large enough to induce significant electrochemicalreaction of interferents, such as urate, ascorbate, and acetaminophen,that may affect the current measurements. Typically the potential isbetween about −150 mV and about +400 mV versus the standard calomelelectrode (SCE). Preferably, the potential of the redox mediator isbetween about −100 mV and +100 mV and, more preferably, the potential isbetween about −50 mV and +50 mV.

The potential may be applied either before or after the sample has beenplaced in the sample chamber. The potential is preferably applied afterthe sample has come to rest in the sample chamber to preventelectrolysis of sample passing through the measurement zone as thesample chamber is filling. When the potential is applied and the sampleis in the measurement zone, an electrical current will flow between theworking electrode and the counter electrode. The current is a result ofthe electrolysis of the analyte in the sample. This electrochemicalreaction occurs via the redox mediator and the optional second electrontransfer agent. For many biomolecules, B, the process is described bythe following reaction equations:

Biochemical B is oxidized to C by redox mediator species A in thepresence of an appropriate enzyme. Then the redox mediator A is oxidizedat the electrode. Electrons are collected by the electrode and theresulting current is measured.

As an example, one sensor of the present invention is based on thereaction of a glucose molecule with two non-leachable ferricyanideanions in the presence of glucose oxidase to produce two non-leachableferrocyanide anions, two protons and gluconolactone. The amount ofglucose present is assayed by electrooxidizing the non-leachableferrocyanide anions to non-leachable ferricyanide anions and measuringthe total charge passed.

Those skilled in the art will recognize that there are many differentreaction mechanisms that will achieve the same result; namely theelectrolysis of an analyte through a reaction pathway incorporating aredox mediator. Equations (1) and (2) are a non-limiting example of sucha reaction.

In a preferred embodiment of the invention, coulometry is used todetermine the concentration of the analyte. This measurement techniqueutilizes current measurements obtained at intervals over the course ofthe assay, to determine analyte concentration. These currentmeasurements are integrated over time to obtain the amount of charge, Q,passed to or from the electrode. Q is then used to calculate theconcentration of the analyte by the following equation:[analyte]=Q/nFV  (3)where n is the number of electron equivalents required to electrolyzethe analyte, F is Faraday's constant (approximately 96,500 coulombs perequivalent), and V is the volume of sample in the measurement zone.

In one embodiment of the invention, the analyte is completely or nearlycompletely electrolyzed. The charge is then calculated from currentmeasurements made during the electrochemical reaction and theconcentration of the analyte is determined using equation (3). Thecompletion of the electrochemical reaction is typically signaled whenthe current reaches a steady-state value. This indicates that all ornearly all of the analyte has been electrolyzed. For this type ofmeasurement, at least 90% of the analyte is typically electrolyzed,preferably, at least 95% of the analyte is electrolyzed and, morepreferably, at least 99% of the analyte is electrolyzed.

For this method it is desirable that the analyte be electrolyzedquickly. The speed of the electrochemical reaction depends on severalfactors, including the potential that is applied between the electrodesand the kinetics of reactions (1) and (2). (Other significant factorsinclude the size of the measurement zone and the presence of sorbent inthe measurement zone.) In general, the larger the potential, the largerthe current through the cell (up to a transport limited maximum) andtherefore, the faster the reaction will typically occur. However, if thepotential is too large, other electrochemical reactions may introducesignificant error in the measurement. Typically, the potential betweenthe electrodes as well as the specific redox mediator and optionalsecond electron transfer agent are chosen so that the analyte will bealmost completely electrolyzed in less than 5 minutes, based on theexpected concentration of the analyte in the sample. Preferably, theanalyte will be almost completely electrolyzed within about 2 minutesand, more preferably, within about 1 minute.

In another embodiment of the invention, the analyte is only partiallyelectrolyzed. The current is measured during the partial reaction andthen extrapolated using mathematical techniques known to those skilledin the art to determine the current curve for the complete or nearlycomplete electrolysis of the analyte. Integration of this curve yieldsthe amount of charge that would be passed if the analyte were completelyor nearly completely electrolyzed and, using equation (3), theconcentration of the analyte is calculated.

The above described methods are based on coulometric analyses, due tothe advantages of coulometric measurements, as described hereinbelow.However, those skilled in the art will recognize that a sensor of theinvention may also utilize potentiometric, amperometric, voltammetric,and other electrochemical techniques to determine the concentration ofan analyte in a sample. There are, however, disadvantages to using someof these techniques. The measurements obtained by these non-coulometricmethods are not temperature independent as the current and potentialobtained by the electrolysis of an analyte on an electrode is verysensitive to sample temperature. This presents a problem for thecalibration of a sensor which will be used to measure bioanalytes andother samples at unknown or variable temperatures.

In addition, the measurements obtained by these non-coulometricelectrochemical techniques are sensitive to the amount of enzymeprovided in the sensor. If the enzyme deactivates or decays over time,the resulting measurements will be affected. This will limit the shelflife of such sensors unless the enzyme is very stable.

Finally, the measurements obtained by non-coulometric electrochemicaltechniques such as amperometry will be negatively affected if asubstantial portion of the analyte is electrolyzed during themeasurement period. An accurate steady-state measurement can not beobtained unless there is sufficient analyte so that only a relativelysmall portion of the analyte is electrolyzed during the measurementprocess.

The electrochemical technique of coulometry overcomes these problems.Coulometry is a method for determining the amount of charge passed orprojected to pass during complete or nearly complete electrolysis of theanalyte. One coulometric technique involves electrolyzing the analyte ona working electrode and measuring the resulting current between theworking electrode and a counter electrode at two or more times duringthe electrolysis. The electrolysis is complete when the current reachesa steady state. The charge used to electrolyze the sample is thencalculated by integrating the measured currents over time. Because thecharge is directly related to the amount of analyte in the sample thereis no temperature dependence of the measurement. In addition, theactivity of the redox mediator does not affect the value of themeasurement, but only the time required to obtain the measurement (i.e.,less active redox mediator requires a longer time to achieve completeelectrolysis of the sample) so that decay of the mediator over time willnot render the analyte concentration determination inaccurate. Andfinally, the depletion of the analyte in the sample by electrolysis isnot a source of error, but rather the objective of the technique.(However, the analyte need not be completely electrolyzed if theelectrolysis curve is extrapolated from the partial electrolysis curvebased on well-known electrochemical principles.)

For coulometry to be an effective measurement technique for determiningthe concentration of an analyte in a sample, it is necessary toaccurately determine the volume of the measured sample. Unfortunately,the volume of the sample in the measurement zone of a small volumesensor (i.e., less than one microliter) may be difficult to accuratelydetermine because the manufacturing tolerances of one or more dimensionsof the measurement zone may have significant variances.

Air-Oxidizable Redox Mediators

Another source of error in a coulometric sensor is the presence ofelectrochemical reactions other than those associated with the analyte.In a sensor having a redox mediator, a potential source of measurementerror is the presence of redox mediator in an unknown mixed oxidationstate (i.e., mediator not reproducibly in a known oxidation state).Redox mediator will then be electrolyzed at the electrode, not inresponse to the presence of an analyte, but simply due to its initialoxidation state. Referring to equations (1) and (2), current notattributable to the oxidation of biochemical B will flow due tooxidation of a portion of a redox mediator, A, that is in its reducedform prior to the addition of the sample. Thus, it is important to knowthe oxidation state of the analyte prior to introduction of the sampleinto the sensor. Furthermore, it is desirable that all or nearly all ofthe redox mediator be in a single oxidation state prior to theintroduction of the sample into the sensor.

Each redox mediator has a reduced form or state and an oxidized form orstate. In one aspect of the invention, it is preferred that the amountof redox mediator in the reduced form prior to the introduction ofsample be significantly smaller than the expected amount of analyte in asample in order to avoid a significant background contribution to themeasured current. In this embodiment of the invention, the molar amountof redox mediator in the reduced form prior to the introduction of theanalyte is preferably less than, on a stoichiometric basis, about 10%,and more preferably less than about 5%, and most preferably less than1%, of the molar amount of analyte for expected analyte concentrations.(The molar amounts of analyte and redox mediator should be comparedbased on the stoichiometry of the applicable redox reaction so that iftwo moles of redox mediator are needed to electrolyze one mole ofanalyte, then the molar amount of redox mediator in the reduced formprior to introduction of the analyte is preferably less than 20% andmore preferably less than about 10% and most preferably less than about2% of the molar amount of analyte for expected analyte concentrations.)Methods for controlling the amount of reduced mediator are discussedbelow.

In another aspect of the invention, it is preferred that the relativeratio of oxidized redox mediator to reduced redox mediator prior tointroduction of the sample in the sensor be relatively constant betweensimilarly constructed sensors. The degree of variation in this ratiobetween similarly constructed sensors will negatively affect the use ofa calibration curve to account for the reduced mediator, as significantvariations between sensors will make the calibration less reliable. Forthis aspect of the invention, the percentage of the redox mediator inthe reduced form prior to introduction of the sample in the sensorvaries by less than about 20% and preferably less than about 10% betweensimilarly constructed sensors.

One method of controlling the amount of reduced redox mediator prior tothe introduction of the sample in the sensor is to provide an oxidizerto oxidize the reduced form of the mediator. One of the most convenientoxidizers is O₂. Oxygen is usually readily available to perform thisoxidizing function. Oxygen can be supplied by exposing the sensor toair. In addition, most polymers and fluids absorb O₂ from the air unlessspecial precautions are taken. Typically, at least 90% of anair-oxidizable (i.e., O₂ oxidizable) mediator is in the oxidized stateupon storage or exposure to air for a useful period of time, e.g., onemonth or less, and preferably, one week or less, and, more preferably,one day or less.

Suitable mediators which are both air-oxidizable (i.e., O₂-oxidizable)and have electron transfer capabilities have been described hereinabove.One particular family of useful mediators are osmium complexes which arecoordinated or bound to ligands with one or more nitrogen-containingheterocycles. In particular, osmium complexed with mono-, di-, andpolyalkoxy-2,2′-bipyridine or mono-, di-, andpolyalkoxy-1,10-phenanthroline, where the alkoxy groups have a carbon tooxygen ratio sufficient to retain solubility in water, areair-oxidizable. These osmium complexes typically have two substitutedbipyridine or substituted phenanthroline ligands, the two ligands notnecessarily being identical. These osmium complexes are furthercomplexed with a polymeric ligand with one or more nitrogen-containingheterocycles, such as pyridine and imidazole. Preferred polymericligands include poly(4-vinyl pyridine) and, more preferably,poly(1-vinyl imidazole) or copolymers thereof.Os[4,4′-dimethoxy-2,2′-bipyridine]₂Cl^(+/+2) complexed with apoly(1-vinyl imidazole) or poly(4-vinyl pyridine) has been shown to beparticularly useful as the Os⁺² cation is oxidizable by O₂ to Os⁺³.Similar results are expected for complexes ofOs[4,7-dimethoxy-1,10-phenanthroline]₂Cl^(+/+2), and other mono-, di-,and polyalkoxy bipyridines and phenanthrolines, with the same polymers.

A complication associated with air-oxidizable mediators arises if theair oxidation of the redox mediator is so fast that a substantialportion of the analyte-reduced redox mediator is oxidized by O₂ duringan analyte assay. This will result in an inaccurate assay as the amountof analyte will be underestimated because the mediator will be oxidizedby the oxidizer rather than by electrooxidation at the electrode. Thus,it is preferred that the reaction of the redox mediator with O₂ proceedsmore slowly than the electrooxidation of the mediator. Typically, lessthan 5%, and preferably less than 1%, of the reduced mediator should beoxidized by the oxidizer during an assay.

The reaction rate of the air oxidation of the mediator can be controlledthrough choice of an appropriate complexing polymer. For example, theoxidation reaction is much faster forOs[4,4′-dimethoxy-2,2′-bipyridine]₂Cl^(+/+2) coordinatively coupled topoly(1-vinyl imidazole) than for the same Os complex coupled topoly(4-vinyl pyridine). The choice of an appropriate polymer will dependon the expected analyte concentration and the potential applied betweenthe electrodes, both of which determine the rate of the electrochemicalreaction.

Thus, in one embodiment of the invention, the preferred redox mediatorhas the following characteristics: 1) the mediator does not react withany molecules in the sample or in the sensor other than the analyte(optionally, via a second electron transfer agent); 2) nearly all of theredox mediator is oxidized by an oxidizer such as O₂ prior tointroduction of the sample in the sensor; and 3) the oxidation of theredox mediator by the oxidizer is slow compared to the electrooxidationof the mediator by the electrode.

Alternatively, if the redox mediator is to be oxidized in the presenceof the analyte and electroreduced at the electrode, a reducer ratherthan an oxidizer would be required. The same considerations for theappropriate choice of reducer and mediator apply as describedhereinabove for the oxidizer.

The use of stable air-oxidizable redox mediators in the electrochemicalsensors of the invention provides an additional advantage during storageand packaging. Sensors of the invention which include air oxidizableredox mediators can be packaged in an atmosphere containing molecularoxygen and stored for long periods of time, e.g., greater than onemonth, while maintaining more than 80% and preferably more than 90% ofthe redox species in the oxidized state.

Optical Sensors

The air-oxidizable redox species of the present invention can be used inother types of sensors. The osmium complexes described hereinabove aresuitable for use in optical sensors, due to the difference in theabsorption spectra and fluorescence characteristics of the complexedOs⁺² and Os⁺³ species. Absorption, transmission, reflection, orfluorescence measurements of the redox species will correlate with theamount of analyte in the sample (after reaction between an analyte andthe redox species, either directly, or via a second electron transferagent such as an enzyme). In this configuration, the molar amount ofredox mediator should be greater, on a stoichiometric basis, than themolar amount of analyte reasonably expected to fill the measurement zoneof the sensor.

Standard optical sensors, including light-guiding optical fiber sensors,and measurement techniques can be adapted for use with theair-oxidizable mediators For example, the optical sensors of theinvention may include a light-transmitting or light reflecting supporton which the air-oxidizable redox species, and preferably ananalyte-responsive enzyme, is coated to form a film. The support filmforms one boundary for the measurement zone in which the sample isplaced. The other boundaries of the measurement zone are determined bythe configuration of the cell. Upon filling the measurement zone with ananalyte-containing sample, reduction of the air-oxidizable mediator bythe analyte, preferably via reaction with the analyte-responsive enzyme,causes a shift in the mediator's oxidation state that is detected by achange in the light transmission, absorption, or reflection spectra orin the fluorescence of the mediator at one or more wavelengths of light.

Multiple Electrode Sensors and Calibration

Errors in assays may occur when mass produced sensor are used because ofvariations in the volume of the measurement zone of the sensors. Two ofthe three dimensions of the measurement zone, the length and the width,are usually relatively large, between about 1-5 mm. Electrodes of suchdimensions can be readily produced with a variance of 2% or less. Thesubmicroliter measurement zone volume requires, however, that the thirddimension be smaller than the length or width by one or two order ofmagnitude. As mentioned hereinabove, the thickness of the sample chamberis typically between about 0.1 and about 0.01 mm. Manufacturingvariances in the thickness may be as large or larger than the desiredthickness. Therefore, it is desirable that a method be provided toaccommodate for this uncertainty in the volume of sample within themeasurement zone.

In one embodiment of the invention, depicted in FIG. 5, multiple workingelectrodes 42, 44, 46 are provided on a base material 48. Theseelectrodes are covered by another base, not shown, which has counterelectrodes, not shown, disposed upon it to provide multiple facingelectrode pairs. The variance in the separation distance between theworking electrode and the counter electrode among the electrode pairs ona given sensor is significantly reduced, because the working electrodesand counter electrodes are each provided on a single base with the samespacer 28 between each electrode pair (see FIG. 3).

One example of a multiple electrode sensor that can be used toaccurately determine the volume of the measurement zones of theelectrode pairs and also useful in reducing noise is presented herein.In this example, one of the working electrodes 42 is prepared with anon-leachable redox mediator and a non-leachable second electrontransfer agent (e.g., an enzyme). Sorbent material may be disposedbetween that working electrode 42 and its corresponding counterelectrode. Another working electrode 44 includes non-leachable redoxmediator, but no second electron transfer agent on the electrode. Again,this second electrode pair may have sorbent material between the workingelectrode 44 and the corresponding counter electrode. An optional thirdworking electrode 46 has no redox mediator and no second electrontransfer agent bound to the electrode, nor is there sorbent materialbetween the working electrode 46 and its corresponding counterelectrode.

The thickness of the sample chamber can be determined by measuring thecapacitance, preferably in the absence of any fluid, between electrode46 (or any of the other electrodes 42, 44 in the absence of sorbentmaterial) and its corresponding counter electrode. The capacitance of anelectrode pair depends on the surface area of the electrodes, theinterelectrode spacing, and the dielectric constant of the materialbetween the plates. The dielectric constant of air is unity whichtypically means that the capacitance of this electrode configuration isa few picofarads (or about 100 picofarads if there is fluid between theelectrode and counter electrode given that the dielectric constant formost biological fluids is approximately 75). Thus, since the surfacearea of the electrodes are known, measurement of the capacitance of theelectrode pair allows for the determination of the thickness of themeasurement zone to within about 1-5%.

The amount of void volume in the sorbent material, can be determined bymeasuring the capacitance between electrode 44 (which has no secondelectron transfer agent) and its associated counter electrode, bothbefore and after fluid is added. Upon adding fluid, the capacitanceincreases markedly since the fluid has a much larger dielectricconstant. Measuring the capacitance both with and without fluid allowsthe determination of the spacing between the electrodes and the voidvolume in the sorbent, and thus the volume of the fluid in the reactionzone.

The sensor error caused by redox mediator in a non-uniform oxidationstate prior to the introduction of the sample can be measured byconcurrently electrolyzing the sample in the measurement zones that areproximate electrodes 42 and 44. At electrode 42, the analyte iselectrolyzed to provide the sample signal. At electrode 44, the analyteis not electrolyzed because of the absence of the second electrontransfer agent (assuming that a second electron transfer agent isnecessary). However, a small charge will pass (and a small current willflow) due to the electrolysis of the redox mediator that was in a mixedoxidation state (i.e., some redox centers in the reduced state and somein the oxidized state) prior to the introduction of the sample. Thesmall charge passed between the electrodes in this second electrode paircan be subtracted from the charge passed between the first electrodepair to substantially remove the error due to the oxidation state of theredox mediator. This procedure also reduces the error associated withother electrolyzed interferents, such as ascorbate, urate, andacetaminophen.

Other electrode configurations can also use these techniques (i.e.,capacitance measurements and coulometric measurements in the absence ofa critical component) to reduce background noise and error due tointerferents and imprecise knowledge of the volume of the interrogatedsample. Protocols involving one or more electrode pairs and one or moreof the measurements described above can be developed and are within thescope of the invention. For example, only one electrode pair is neededfor the capacitance measurements, however, additional electrode pairsmay be used for convenience.

EXAMPLES

The invention will be further characterized by the following examples.These examples are not meant to limit the scope of the invention whichhas been fully set forth in the foregoing description. Variations withinthe concepts of the invention are apparent to those skilled in the art.

Example 1 Preparation of a Small Volume In Vitro Sensor for theDetermination of Glucose Concentration

A sensor was constructed corresponding to the embodiment of theinvention depicted in FIG. 1. The working electrode was constructed on aMylar™ film (DuPont), the Mylar™ film having a thickness of 0.175 mm anda diameter of about 2.5 cm. An approximately 12 micron thick carbon padhaving a diameter of about 1 cm was screen printed on the Mylar film.The carbon electrode was overlaid with a water-insoluble dielectricinsulator (Insulayer) having a thickness of 12 μm, and a 4 mm diameteropening in the center.

The center of the carbon electrode, which was not covered by thedielectric, was coated with a redox mediator. The redox mediator wasformed by complexing poly(1-vinyl imidazole) withOs(4,4′-dimethoxy-2,2′-bipyridine)₂Cl₂ followed by cross-linking glucoseoxidase with the osmium polymer using polyethylene glycol diglycidylether as described in Taylor, et al., J. Electroanal. Chem., 396:511(1995). The ratio of osmium to imidazole functionalities in the redoxmediator was approximately 1:15. The mediator was deposited on theworking electrode in a layer having a thickness of 0.6 μm and a diameterof 4 mm. The coverage of the mediator on the electrode was about 60μg/cm (dry weight). A spacer material was placed on the electrodesurrounding the mediator-covered surface of the electrode. The spacerwas made of poly(tetrafluoroethylene) (PTFE) and had a thickness ofabout 0.040 mm.

A sorbent material was placed in contact with the mediator-coveredsurface of the working electrode. The sorbent was made of nylon (TetkoNitex nylon 3-10/2) and had a diameter of 5 mm, a thickness of 0.045 mm,and a void volume of about 20%. The volume of sample in the measurementzone was calculated from the dimensions and characteristics of thesorbent and the electrode. The measurement zone had a diameter of 4 mm(the diameter of the mediator covered surface of the electrode) and athickness of 0.045 mm (thickness of the nylon sorbent) to give a volumeof 0.57 μL. Of this space, about 80% was filled with nylon and the other20% was void space within the nylon sorbent. This resulting volume ofsample within the measurement zone was about 0.11 μL.

A counter/reference electrode was placed in contact with the spacer andthe side of the sorbent opposite to the working electrode so that thetwo electrodes were facing each other. The counter/reference electrodewas constructed on a Mylar™ film having a thickness of 0.175 mm and adiameter of about 2.5 cm onto which a 12 micron thick layer ofsilver/silver chloride having a diameter of about 1 cm was screenprinted.

The electrodes, sorbent, and spacer were pressed together using plateson either side of the electrode assembly. The plates were formed ofpolycarbonate plastic and were securely clamped to keep the sensortogether. The electrodes were stored in air for 48 hours prior to use.

Tabs extended from both the working electrode and the counter/referenceelectrode and provided for an electrical contact with the analyzingequipment. A potentiostat was used to apply a potential difference of+200 mV between the working and counter/reference electrodes, with theworking electrode being the anode. There was no current flow between theelectrodes in the absence of sample, which was expected, as noconductive path between the electrodes was present.

The sample was introduced via a small tab of nylon sorbent materialformed as an extension from the nylon sorbent in the sample chamber.Liquid was wicked into the sorbent when contact was made between thesample and the sorbent tab. As the sample chamber filled and the samplemade contact with the electrodes, current flowed between the electrodes.When glucose molecules in the sample came in contact with the glucoseoxidase on the working electrode, the glucose molecules wereelectrooxidized to gluconolactone. The osmium redox centers in the redoxmediator then reoxidized the glucose oxidase. The osmium centers were inturn reoxidized by reaction with the working electrode. This provided acurrent which was measured and simultaneously integrated by acoulometer. (EG&G Princeton Applied Research Model #173)

The electrochemical reaction continued until the current reached asteady state value which indicated that greater than 95% of the glucosehad been electroreduced. The current curve obtained by measurement ofthe current at specific intervals was integrated to determine the amountof charge passed during the electrochemical reaction. These charges werethen plotted versus the known glucose concentration to produce acalibration curve.

The sensor was tested using 0.5 μL aliquots of solutions containingknown concentrations of glucose in a buffer of artificial cerebrospinalfluid or in a control serum (Baxter-Dade, Monitrol Level 1, Miami, Fla.)in the range of 3 to 20 mM glucose. The artificial cerebrospinal fluidwas prepared as a mixture of the following salts: 126 mM NaCl, 27.5 mMNaHCO₃, 2.4 mM KCl, 0.5 mM KH₂PO₄, 1.1 mM CaCl₂-2H₂O, and 0.5 mM Na₂SO₄.

The results of the analyses are shown in Table 1 and in FIG. 7. In Table1, Q_(avg) is the average charge used to electrolyze the glucose in 3-6identical test samples (FIG. 7 graphs the charge for each of the testsamples) and the 90% rise time corresponds to the amount of timerequired for 90% of the glucose to be electrolyzed. The data show asensor precision of 10-20%, indicating adequate sensitivity of thesensor for low glucose concentrations, as well as in the physiologicallyrelevant range (30 μg/dL-600 μg/dL).

TABLE 1 Sensor Results Using Glucose Oxidase Number of Samples 90% risetime Tested Q_(avg) (μC) (sec) buffer only 4  9.9 ± 1.8 13 ± 6 3 mMglucose/buffer 5 17.8 ± 3.5 19 ± 5 6 mM glucose/buffer 4 49.4 ± 4.9 25 ±3 10 mM glucose/buffer 6  96.1 ± 12.4  36 ± 17 15 mM glucose/buffer 5205.2 ± 75.7  56 ± 23 20 mM glucose/buffer 4 255.7 ± 41.0  62 ± 17 4.2mM glucose/serum 3 44.2 ± 4.3 44 ± 3 15.8 mM glucose/serum 3 218.2 ±57.5  72 ± 21

The average measured values of glucose concentration were fit by one ormore equations to provide a calibration curve. FIG. 8 shows thecalibration curves for the glucose/buffer data of Table 1. One of the15.0 mM glucose measurements was omitted from these calculations becauseit was more than two standard deviations away from the average of themeasurements. The higher glucose concentrations (10-20 mM) were fit by alinear equation. The lower glucose concentrations were fit by a secondorder polynomial.

FIG. 9 shows the data of Table 1 plotted on an error grid developed byClarke, et al. Diabetes Care, 5, 622-27, 1987, for the determination ofthe outcome of errors based on inaccurate glucose concentrationdetermination. The graph plots “true” glucose concentration vs. measuredglucose concentration, where the measured glucose concentration isdetermined by calculating a glucose concentration using the calibrationcurves of FIG. 8 for each data point of FIG. 7. Points in zone A areaccurate, those in zone B are clinically acceptable, and those in zonesC, D, and E lead to increasingly inappropriate and finally dangeroustreatments.

There were 34 data points. Of those data points 91% fell in zone A, 6%in zone B, and 3% in zone C. Only one reading was determined to be inzone C. This reading was off-scale and is not shown in FIG. 9. Thus, 97%of the readings fell in the clinically acceptable zones A and B.

The total number of Os atoms was determined by reducing all of the Osand then electrooxidizing it with a glucose-free buffer in the samplechamber. This resulted in a charge of 59.6±5.4 μC. Comparison of thisresult with the glucose-free buffer result in Table 1 indicated thatless than 20% of the Os is in the reduced form prior to introduction ofthe sample. The variability in the quantity of osmium in the reducedstate is less than 5% of the total quantity of osmium present.

Example 2 Response of the Glucose Sensor to Interferents

A sensor constructed in the same manner as described above for Example 1was used to determine the sensor's response to interferents. The primaryelectrochemical interferents for blood glucose measurements areascorbate, acetaminophen, and urate. The normal physiological ortherapeutic (in the case of acetaminophen) concentration ranges of thesecommon interferents are:

ascorbate: 0.034-0.114 mM

acetaminophen: 0.066-0.200 mM

urate (adult male): 0.27-0.47 mM

Tietz, in: Textbook of Clinical Chemistry, C. A. Burtis and E. R.Ashwood, eds., W.B. Saunders Co., Philadelphia 1994, pp. 2210-12.

Buffered glucose-free interferent solutions were tested withconcentrations of the interferents at the high end of the physiologicalor therapeutic ranges listed above. The injected sample volume in eachcase was 0.5 μL. A potential of +100 mV or +200 mV was applied betweenthe electrodes. The average charge (Q_(avg)) was calculated bysubtracting an average background current obtained from a buffer-only(i.e., interferent-free) solution from an average signal recorded withinterferents present. The resulting average charge was compared with thesignals from Table 1 for 4 mM and 10 mM glucose concentrations todetermine the percent error that would result from the interferent.

TABLE 2 Interferent Response of Glucose Sensors Error @ Error @ 4 mM 10mM Solution E (mV) N Q_(avg) (μC) glucose glucose 0.114 mM ascorbate 1004 0.4 2% <1% 0.114 mM ascorbate 200 4 −0.5 2% <1%  0.2 mM acetaminophen100 4 0.1 <1%   <1%  0.2 mM acetaminophen 200 4 1.0 5%   1%  0.47 mMurate 100 4 6.0 30%    7%  0.47 mM urate 200 4 18.0 90%  21%

These results indicated that ascorbate and acetaminophen were notsignificant interferents for the glucose sensor configuration,especially for low potential measurements. However, urate providedsignificant interference. This interference can be minimized bycalibrating the sensor response to a urate concentration of 0.37 mM,e.g., by subtracting an appropriate amount of charge as determined byextrapolation from these results from all glucose measurements of thesensor. The resulting error due to a 0.10 mM variation in urateconcentration (the range of urate concentration is 0.27-0.47 in an adultmale) would be about 6% at 4 mM glucose and 100 mV.

Example 3 Sensor with Glucose Dehydrogenase

A sensor similar to that described for Example 1 was prepared and usedfor this example, except that glucose oxidase was replaced bypyrroloquinoline quinone glucose dehydrogenase and a potential of only+100 mV was applied as opposed to the +200 mV potential in Example 1.The results are presented in Table 3 below and graphed in FIG. 10.

TABLE 3 Sensor Results Using Glucose Dehydrogenase n Q_(avg) (μC) 90%rise time(s) buffer 4 21.7 ± 5.2 14 ± 3 3 mM glucose/buffer 4  96.9 ±15.0 24 ± 6 6 mM glucose/buffer 4 190.6 ± 18.4 26 ± 6 10 mMglucose/buffer  6 327.8 ± 69.3 42 ± 9

The results indicated that the charge obtained from the glucosedehydrogenase sensor was much larger than for the comparable glucoseoxidase sensor, especially for low concentrations of glucose. For 4 mMglucose concentrations the measurements obtained by the two sensorsdiffered by a factor of five. In addition, the glucose dehydrogenasesensor operated at a lower potential, thereby reducing the effects ofinterferent reactions.

In addition, the results from Table 3 were all fit by a linearcalibration curve as opposed to the results in Example 1, as shown inFIG. 10. A single linear calibration curve is greatly preferred tosimplify sensor construction and operation.

Also, assuming that the interferent results from Table 2 are applicablefor this sensor, all of the interferents would introduce an error ofless than 7% for a 3 mM glucose solution at a potential of 100 mV.

Example 4 Determination of Lactate Concentration in a Fluid Stream

The sensor of this Example was constructed using a flow cell(BioAnalytical Systems, Inc. # MF-1025) with a glassy carbon electrode.A redox mediator was coated on the electrode of the flow cell to providea working electrode. In this case, the redox mediator was a polymerformed by complexing poly(1-vinyl imidazole) withOs(4,4′-dimethyl-2,2′-bipyridine)₂Cl₂ with a ratio of 1 osmium for every15 imidazole functionalities. Lactate oxidase was cross-linked with thepolymer via polyethylene glycol diglycidyl ether. The mediator wascoated onto the electrode with a coverage of 500 μg/cm² and a thicknessof 5 μm. The mediator was covered by a polycarbonate track-etchedmembrane. (Osmonics-Poretics #10550) to improve adherence in the flowstream. The membrane was then overlaid by a single 50 μm thick spacergasket (BioAnalytical Systems, Inc. #MF-1062) containing a void whichdefined the sample chamber and corresponding measurement zone. Assemblyof the sensor was completed by attachment of a cell block (BioAnalyticalSystems, Inc. #MF-1005) containing the reference and auxiliaryelectrodes of the flow cell.

The sample chamber in this case corresponded to a 50 μm thick cylinder(the thickness of the spacer gasket) in contact with a mediator-coatedelectrode having a surface area of 0.031 cm². The calculated volume ofsample in the measurement zone of this sensor was approximately 0.16 μL.

The flow rate of the fluid stream was 5 μL/min. A standard threeelectrode potentiostat was attached to the cell leads and a potential of+200 mV was applied between the redox mediator-coated glassy carbonelectrode and the reference electrode. This potential was sufficient todrive the enzyme-mediated oxidation of lactate.

As the fluid stream flowed through the sensor, a steady-state currentproportional to the lactate concentration was measured. At periodicintervals the fluid flow was stopped and current was allowed to flowbetween the electrodes until approximately all of the lactate in themeasurement zone was electrooxidized, as indicated by the achievement ofa stabilized, steady-state current. The total charge, Q, required forlactate electrooxidation was found by integration of the differentialcurrent registered from the flow stoppage until the current reached asteady-state. The concentration was then calculated by the followingequation:[lactate]=Q/2FV  (4)

where V is the volume of sample within the measurement zone and F isFaraday's constant.

This assay was performed using lactate solutions having nominal lactateconcentrations of 1.0, 5.0, and 10.0 mM. The measured concentrations forthe assay were 1.9, 5.4, and 8.9 mM respectively.

Example 5 Determination of the Oxidation State ofOs(4,4′-dimethoxy-2,2′-bipyridine)₂Cl^(+/+2) Complexed with poly(1-vinylimidazole)

A sensor having a three electrode design was commercially obtained fromEcossensors Ltd., Long Hanborough, England, under the model name “largearea disposable electrode”. The sensor contained parallel and coplanarworking, reference and counter electrodes. The working surface area (0.2cm²) and counter electrodes were formed of printed carbon and thereference electrode was formed of printed Ag/AgCl. A redox mediator wascoated on the carbon working electrode. The redox mediator was formed bycomplexation of poly(1-vinyl imidazole) withOs(4,4′-dimethoxy-2,2′-bipyridine)₂Cl₂ in a ratio of 15 imidazole groupsper Os cation followed by crosslinking the osmium polymer with glucoseoxidase using polyethylene glycol diglycidyl ether.

The electrode was cured at room temperature for 24 hours. The coplanarelectrode array was then immersed in a buffered electrolyte solution,and a potential of +200 mV (sufficient for conversion of Os(II) toOs(III),) was applied between the working electrode and the referenceelectrode.

Upon application of the potential, an undetectable charge of less than 1μC was passed. Subsequent reduction and reoxidation of the redoxmediator yielded a charge for conversion of all Os from Os(II) toOs(III) of 65 μC. Therefore, more than 98% of the Os cations in theredox mediator were in the desired oxidized Os(III) state.

Example 6 Determination of the Oxidation State of theOs(4,4′-dimethoxy-2,2′-bipyridine₂Cl^(+/+2) Complexed with poly(4-vinylpyridine)

A similar experiment to that of Example 5 was conducted with the sameworking/counter/reference electrode configuration except that the redoxmediator on the working electrode was changed to a complex ofOs(4,4′-dimethoxy-2,2′-bipyridine)₂Cl₂ with poly(4-vinyl pyridine), with12 pyridine groups per Os cation, cross linked with glucose oxidase viapolyethylene glycol diglycidyl ether.

Two sensors were constructed. The electrodes of the two sensors werecured at room temperature for 24 hours. The electrodes were thenimmersed in a buffered electrolyte solution and a potential of +200 mVwas applied between the working and reference electrodes.

Upon application of the potential to the electrodes, a charge of 2.5 μCand 3.8 μC was passed in the two sensors, respectively. Subsequentreduction and reoxidation of the redox mediators yielded oxidationcharges of 27.9 μC and 28.0 μC, respectively. Therefore, the sensorsoriginally contained 91% and 86% of the Os cations in the desirableoxidized Os(III) state.

Example 7 Optical Sensor

An optical sensor is constructed by applying a film of redox polymerwith crosslinked enzyme onto a light-transparent support such as a glassslide. The quantity of redox mediator is equal to or greater than (in astoichiometric sense) the maximum quantity of analyte expected to fillthe measurement zone. The spacer material, sorbent and facing supportare securely clamped. The sample chamber is adapted to transmit lightthrough the assembled sensor to an optical density detector or to afluorescence detector. As sample fills the sample chamber and the redoxmediator is oxidized, changes in the absorption, transmission,reflection or fluorescence of the redox mediator in the chamber arecorrelated to the amount of glucose in the sample.

Example 8 Blood Volumes from Upper Arm Lancet Sticks

The forearm of a single individual was pierced with a lancet multipletimes in order to determine the reproducibility of blood volumesobtained by this method. Despite more than thirty lancet sticks in theanterior portion of each forearm and the dorsal region of the leftforearm, the individual identified each stick as virtually painless.

The forearm was pierced with a Payless Color Lancet. The blood from eachstick was collected using a 1 μL capillary tube, and the volume wasdetermined by measuring the length of the blood column. The volumesobtained from each stick are shown in Table 4.

TABLE 4 Volume of Lancet Sticks Left Anterior Right Anterior Left DorsalForearm, (nL) Forearm, (nL) Forearm, (nL) 1 180 190 180 2 250 180 300 3170 120 310 4 150 100 300 5 100 210 60 6 50 140 380 7 90 120 220 8 130140 200 9 120 100 380 10  100 320 11  260 12  250 13  280 14  260 Avg.138 ± 58 nL 140 ± 40 nL 264 ± 83 nL

The invention has been described with reference to various specific andpreferred embodiment and techniques. However, it will be apparent to oneof ordinarily skill in the art that many variation and modifications maybe made while remaining within the spirit and scope of the invention.

All publications and patent applications in this specification areindicative of the level of ordinary skill in the art to which thisinvention pertains. All publications and patent applications are hereinincorporated by reference to the same extent as if each individualpublication or patent application was specifically and individuallyindicated by reference.

1. An analyte measurement device for determining the concentration ofglucose in blood from a patient, comprising: a sensor comprising: aworking electrode, a counter/reference electrode, a redox mediator, aglucose-responsive enzyme, and a polymer, wherein said workingelectrode, counter/reference electrode, redox mediator, enzyme andpolymer are in communication with a sample chamber of the sensor havinga volume of less than about 0.5 μl; and a tapered tip comprising asample application site; and a coulometer.
 2. The sensor of claim 1,wherein the sensor is in the shape of a strip.
 3. The sensor of claim 1,wherein the sensor includes a sorbent material for transporting blood tocontact the working and counter electrodes.
 4. The sensor of claim 3,wherein the sorbent material is contained in the sample chamber.
 5. Thesensor of claim 1, wherein the glucose-responsive enzyme comprises aglucose oxidase or a glucose dehydrogenase.
 6. The sensor of claim 1,wherein the working electrode is on a first substrate and the counterelectrode is on a second substrate.
 7. The sensor of claim 1, whereinthe polymer is poly(4-vinyl pyridine) (PVP) or poly(l-vinyl imidazole)(PVI).
 8. The sensor of claim 1, wherein the sensor further comprises athird electrode.
 9. The sensor of claim 1, wherein the charge generatedby the sensor by electrolysis of a buffer solution comprising no glucoseis no more than about 26% of the charge generated by electrolysis of abuffer solution comprising 6 mM glucose.
 10. The sensor of claim 1,wherein the charge generated by the sensor by electrolysis of a buffersolution comprising no glucose is no more than about 14% of the chargegenerated by electrolysis of a buffer solution comprising 10 mM glucose.11. The sensor of claim 1, wherein the redox mediator comprises anosmium complex, a ferrocyanide, or a ferricyanide.
 12. The sensor ofclaim 1, wherein the working electrode is constructed of a materialselected from gold, carbon, platinum, ruthenium dioxide and palladium.13. The sensor of claim 1, wherein the working electrode is separatedfrom the counter/reference electrode by a distance of less than 200micrometer.
 14. A method for determining the blood glucose concentrationin a patient, the method comprising: (a) providing a sensor comprising:a working electrode, a counter/reference electrode, a redox mediator, aglucose-responsive enzyme, and a polymer, wherein said workingelectrode, counter/reference electrode, redox mediator, enzyme andpolymer are in communication with a sample chamber of the sensor havinga volume of less than about 0.5 μl; and a tapered tip comprising asample application site; and (b) determining the blood glucoseconcentration by coulometry.
 15. The method of claim 14, wherein themethod comprises contacting the patient with the sensor and thendetermining the blood glucose concentration.
 16. The method of claim 14,wherein the charge generated by the sensor by electrolysis of a buffersolution comprising no glucose is no more than about 26% of the chargegenerated by electrolysis of a buffer solution comprising 6 mM glucose.17. The sensor of claim 14, wherein the charge generated by the sensorby electrolysis of a buffer solution comprising no glucose is no morethan about 14% of the charge generated by electrolysis of a buffersolution comprising 10 mM glucose.
 18. An analyte measurement device fordetermining the concentration of glucose in blood from a patient,comprising: a sensor comprising: a working electrode, acounter/reference electrode, a redox mediator, a glucose-responsiveenzyme and a polymer, wherein said working electrode is on a firstsubstrate and said counter/reference electrode is on a second substrate,and said working electrode, counter/reference electrode, redox mediator,enzyme and polymer are in communication with a sample chamber of thesensor having a volume of less than about 0.5 μl; and a tapered tipcomprising a sample application site; and a coulometer.
 19. The sensorof claim 18, wherein the charge generated by the sensor by electrolysisof a buffer solution comprising no glucose is no more than about 26% ofthe charge generated by electrolysis of a buffer solution comprising 6mM glucose.
 20. The sensor of claim 18, wherein the charge generated bythe sensor by electrolysis of a buffer solution comprising no glucose isno more than about 14% of the charge generated by electrolysis of abuffer solution comprising 10 mM glucose.
 21. The sensor of claim 18,wherein the glucose-enzyme comprises glucose oxidase or glucosedehydrogenase.
 22. The sensor of claim 18, wherein the redox mediatorcomprises an osmium complex, a ferrocyanide, or a ferricyanide.
 23. Thesensor of claim 18, wherein the working electrode is constructed of amaterial selected from gold, carbon, platinum, ruthenium dioxide andpalladium.
 24. The sensor of claim 18, wherein the working electrode isseparated from the counter/reference electrode by a distance of lessthan 200 micrometer.